Use and making of biosensors utilizing antimicrobial peptides for highly sensitive biological monitoring

ABSTRACT

A biosensor and method of making are disclosed. The biosensor is configured to detect a target and may include a peptide immobilized on a sensing component, the sensing component having an anode and a cathode. The immobilized peptide may comprise an antimicrobial peptide binding motif for the target. The sensing component has an electrical conductivity that changes in response to binding of the immobilized peptide to the target. The immobilized peptide may bind one or more targets selected from the list consisting of: bacteria, Gram-negative bacteria, Gram-positive bacteria, pathogens, protozoa, fungi, viruses, and cancerous cells. The biosensor may have a display with a readout that is responsive to changes in electrical conductivity of the sensing component. The display unit may be wirelessly coupled to the sensing component. A resonant circuit with an inductive coil may be electrically coupled to the sensing component. A planar coil antenna may be disposed in proximity to the resonant circuit, the planar coil antenna being configured to provide power to the sensing component.

CROSS-REFERENCE TO PRIOR FILED APPLICATION

This application is a continuation application of, and claims priorityto, an earlier filed nonprovisional U.S. patent application Ser. No.13/171,120 filed on Jun. 28, 2011 which is U.S. Pat. No. 9,029,168issuing May 12, 2015, and also claims priority to an earlier filed U.S.provisional application 61/359,070 filed on Jun. 28, 2010, both of whichare herein incorporated by reference in their entirety.

UNITED STATES GOVERNMENT RIGHTS

This invention was made with government support under Air Force Officeof Scientific Research via a Young Investigator Grant(#FA9550-09-1-0096) and as a fellow of the American Asthma Foundation(09-0038). The government has certain rights to this invention.

FIELD OF INVENTION

This invention relates to the field of biosensors; more specifically theinvention related to the detection of Gram-negative, Gram-positivebacteria and other biologic targets through biosensors utilizingimmobilized antimicrobial peptides. Certain embodiments of thisinvention also relate to the field of nanotechnology.

BACKGROUND

Bacterial infections remain the leading cause of death in developingnations, accounting for an estimated 40% of deaths (Ivnitski D et al.(1999) Biosens Bioelectron 14:599-624.). For instance, the strainO157:H7 of E. coli is considered to be one of the most dangerous foodborne pathogens (Buchanan R L, et al. (1997) Food Technol 51:69-76 andJay J M ed. (1992) in Modern food microbiology. (Van Nostrand Reinhold,N.Y.)). In developed countries, bacterial contamination is also ofcritical concern, particularly in the pharmaceutical industry, where themost reliable test for contamination is the detection of endotoxins withhorseshoe crab blood (Walls E A, Berkson J, Smith S A (2002) Rev FishSci 10:39-73). Microbial infections and drug-resistant supergerms arealso a leading cause of military deaths, particularly in soldiers withburn injuries, and are considered potential biowarfare agents (Compton JA F (1987) in Military chemical and biological agents: Chemical andtoxicological properties. (Telford press, Caldwell, N.J.) pp: 458,Malcolm D ed. (1994) in Biological Warfare in the 21st Century.(Brassey's, UK):258, and D'Avignon L C, et al (Jan. 13, 2010)Contribution of bacterial and viral infections to attributable mortalityin patients with severe burns: An autopsy series. Burns).

While containment strategies such as vaccination and “broadband”antibiotic usage in hospitals have helped reduce the severity ofbacterial infections, these strategies have also inadvertently promotedthe emergence of antibiotic resistance. Thus, the development of asensor that can detect the presence of an infectious outbreak from abroad spectrum of pathogenic species would be desirable. It would beespecially desirable if the sensor was scalable so that it could be usedin nearly every situation where monitoring is desired, including thelaboratory, in the field at the source of potential infection, and on orinside animals, including humans, that many be exposed to infection.

SUMMARY OF THE INVENTION

A biosensor and method of making are disclosed. The biosensor isconfigured to detect a target and may include a peptide immobilized on asensing component, the sensing component having an anode and a cathode.The immobilized peptide may comprise an antimicrobial peptide bindingmotif for the target. The sensing component has an electricalconductivity that changes in response to binding of the immobilizedpeptide to the target. The target may be a bacterium, a Gram-negativebacterium, Escherichia coli and/or Escherichia coli O157:H7. Theimmobilized peptide may bind one or more targets selected from the listconsisting of: bacteria, Gram-negative bacteria, Gram-positive bacteria,pathogens, protozoa, fungi, viruses, and cancerous cells.

The biosensor may have a display unit electrically coupled to thesensing component forming an electric circuit, the display unit having areadout that is responsive to changes in electrical conductivity of thesensing component. The display unit may be wirelessly coupled to thesensing component, the display unit having a readout that is responsiveto changes in electrical conductivity of the sensing component. A powersource may be electrically coupled to the sensing component. A wirelesstelemetry device may be electrically coupled to the sensing component. Aresonant circuit with an inductive coil may be electrically coupled tothe sensing component. A planar coil antenna may be disposed inproximity to the resonant circuit, the planar coil antenna beingconfigured to provide power to the sensing component.

The sensing component may have a longest linear dimension of less than 1mm. The sensing component may have a longest linear dimension of lessthan 1 μm. The display unit may use impedance spectroscopy to determinethe electrical conductivity of the electric circuit and change thereadout accordingly. The sensing component may be an interdigitatedmicroelectrode array. The sensing component may be a graphene substrate.The sensing component may be a nanowire substrate.

A method of using an immobilized antimicrobial peptide to detect atarget is disclosed, the method may include exposing the sample to atleast one peptide immobilized on a sensing component, the immobilizedpeptide having an antimicrobial binding motif, the electricalconductivity of the sensing component changing in response to binding ofthe immobilized peptide with a target. The electrical conductivity ofthe sensing component is measured. The presence or absence of the targetin the sample is determined based on the electrical conductivity of thesensing component. An electric circuit may be formed including thesensing component and a power source. The electric circuit may bepowered with a planar coil antenna disposed in proximity to the electriccircuit. A concentration of a target may be determined based on theelectrical conductivity of the sensing component.

A method of making a biosensor configured to detect a target isdisclosed. The method may include selecting an antimicrobial peptidebased on its Polyvinyl binding affinity for the target. Theantimicrobial peptide may be immobilized on a sensing component, thesensing component having an electrical conductivity that changes basedon binding to the target. An anode and a cathode are coupled to thesensing component. The antimicrobial peptide may be modified prior tobeing immobilized. The biosensor may be embedded on a bioresorbablesubstrate. The bioresorbable substrate may be selected from a listconsisting of: poly(ethylene terepthalate), poly(imide), poly(ethersulfone), cellulose, paper, silk and silk fibroin.

The biosensor may include an antimicrobial peptide immobilized on amicrobe, the microprobe having an electrical conductivity that isdependent on a binding state of the binding molecule. An anode and acathode may be coupled to the microprobe. A display with a readout maybe electrically coupled to the cathode and anode, forming a circuit, thereadout being based on the electrical conductivity of the microprobe.

The biosensor may include an interdigitated microelectrode arraydisposed on a bioresorbable substrate. An anode and a cathode may beelectrically coupled to the interdigitated microelectrode array. Aresonant circuit with an inductive coil may be electrically coupled tothe cathode and anode. A wireless telemetry unit may be electricallycoupled to the resonant circuit, forming an electric circuit. A bindingmolecule may be disposed on the interdigitated microelectrode array, theinterdigitated microelectrode having an electrical conductivity that isdependent on a binding state of the binding molecule. The bindingmolecule may be selected from the list consisting of: antimicrobialpeptides, antibodies, modified antimicrobial peptides, modifiedantibodies, chimeric peptides containing antimicrobial peptide bindingmotifs, chimeric peptides containing antibody binding motifs, DNAfragments, RNA fragments, peptide binding motifs, proteins, smallmolecules and polymers combinations thereof. The bioresorbable substrateis selected from a list consisting of: poly(ethylene terepthalate),poly(imide), poly(ether sulfone), cellulose, paper, silk, silk fibroinand combinations thereof.

The bioresorbable substrate may be placed in contact with a biologicsurface. The biologic surface may be capable of bioresorption. Thebiologic surface may be selected from the list consisting of: teeth,bone, skin, tissue, hair, nail, cornea, gum, tongue, palate, brain,heart, lung, membrane, leaf, root, bark, fur, feather, chiton and scale.The biosensor may also include a receiver unit having a planar coilantenna electrically coupled to a readout, the planar coil antenna beingconfigured to power the electric circuit, enabling the wirelesstelemetry unit to send a wireless signal corresponding to an electricalconductivity of the circuit to the receiver unit, the readout beingresponsive to changes in the electrical conductivity of the circuit. Thebinding molecule may contain a glucose-binding motif.

BRIEF DESCRIPTION OF THE FIGURES

FIG. 1A depicts a schematic of the impedance spectroscopy measurementsetup;

FIG. 1B depicts a simplified equivalent circuit of the microelectrodearray/electrolyte interface before bacterial binding;

FIG. 1C depicts a simplified equivalent circuit of the microelectrodearray/electrolyte interface after bacterial binding;

FIG. 2A depicts an artistic illustration of AMPs immobilized on aninterdigitated microelectrode array;

FIG. 2B depicts a magnified artistic illustration image of the AMPmagainin I in helical form, modified with a terminal cysteine residue,and with clearly defined hydrophobic and hydrophilic faces;

FIG. 2C depicts an artistic illustration of the detection of bacteriavia binding of target cells to the immobilized AMPs;

FIG. 2D depicts an optical image of the interdigitated microelectrodearray (scale bar: 50 μm);

FIG. 3A depicts a line graph of the sensitivity of the AMP electronicbiosensor of impedance spectra of various concentrations of E. coliO157:H7 cells (lines with geometric shapes), of a nonlabeled sensor(solid line), and of a sensor with an N-terminal immobilized AMP (dashedline);

FIG. 3B depicts a bar graph of the impedance spectra of variousconcentrations of E. coli with the AMP sensor at 10 Hz, error bars showstandard deviation (N=3);

FIG. 4 depicts a line graph of the impedance spectra of variousconcentrations of E. coli O157:H7 cells after exposure to a nonlabeled“blank” sensor, the inset shows the optical micro-graph of the baresensor after exposure to E. coli cells of concentration 10⁷ cfu/mL;

FIG. 5 depicts a line graph of the effect of the surface density ofimmobilized magainin I on the binding of bacterial cells, error barsshow standard deviation, (N=3); the left inset of FIG. 5 depicts anillustration antimicrobial peptide (Magainin I) immobilized on to goldsubstrate; the right inset of FIG. 5 depicts a fluorescent image ofantimicrobial peptide (Magainin I) immobilized on to gold substrateafter labeling with FITC dye;

FIGS. 6A-C depict the optical microscopy of the selectivity of AMPs, theleft panels are pictures of selective binding of the immobilized AMP tovarious stained bacterial cells (10⁷ cfu/mL), including: FIG. 6A E. coliO157:H7, FIG. 6B S. typhimurium, FIG. 6C E. coli ATCC 35218, and FIG. 6DL. monocytogenes, the right panels are graphs of the correspondingsurface density of the bound cell, scale bars are 10 μm;

FIG. 7A depicts a line graph of the impedance spectra of theAMP-functionalized microelectrode array after interaction with variousbacterial samples (10⁷ cfu/mL);

FIG. 7B depicts a bar graph of the impedance changes associated withvarious bacterial species at 10 Hz, error bars show standard deviation(N=3);

FIG. 8A shows a photograph of the microfluidic flow cell duringreal-time binding of bacteria to AMP biosensors;

FIG. 8B shows an optical micrograph of the microfluidic channel with anembedded interdigitated microelectrode array chip;

FIG. 8C shows an optical image of the embedded microelectrode arrayafter exposure to 10⁷ cfu/mL bacterial cells for 30 min;

FIG. 8D depicts a graph of real-time monitoring of the interaction ofthe AMP-functionalized sensor (and an unlabeled control chip) withvarious concentrations of E. coli cells;

FIG. 9A shows the optical image of microfluidic functionalizationchannels (vertical conduits) intersecting nanowire sensor devices; thenanowire islands (horizontal bars) are electrically contacted by metalleads; the inset depicts a scanning electron micrograph of the nanowirefilm;

FIG. 9B depicts the characterization of the bare silicon-on-insulator(SOI), amine-terminated (APTES), and peptide-coupled surfaces by X-rayphotoelectron spectroscopy, the inset depicts water contact anglegoniometric measurements of the surfaces. FIGS. 9A-F depicts an artisticillustration of the steps for transfer printing SNAP nanowires ontoplastic substrates;

FIG. 10A depicts an artistic illustration of first step in which thenanowires are etched into a single-crystal silicon-on-insulatorsubstrate;

FIG. 10B depicts an artistic illustration of the step in which theexposed oxide is etched and a piece of PDMS makes conformal contact withthe nanowire surfaces;

FIG. 10C depicts an artistic illustration of the step in which the PDMSwith adhered nanowires is peeled back from the host substrate;

FIG. 10D depicts an artistic illustration of the step in which theplastic substrate is spin-cast with epoxy;

FIG. 10E depicts an artistic illustration of the step in which the PDMSmakes conformal contact with the plastic, and the epoxy is cured;

FIG. 10F depicts an artistic illustration of final step in which peelingback the PDMS leaves behind the SNAP nanowires in their originalorientation, but on plastic.

FIGS. 11A-D depict artistic illustrations of a graphene/silk-basedbioresorbable passive wireless biosensor;

FIG. 11A depicts an artistic illustration of a graphene based wirelessbiosensor with a hybrid sensing element and a passive wireless telemetrysystem on a bioresorbable platform of silk fibroin;

FIG. 11B depicts an artistic illustration of the biosensor afterbioresorption of the silk fibroin and the attachment of the biosensor onto the surface of a tooth;

FIG. 11C depicts a magnified artistic illustration of the bioresorbedsensing element, illustrating the capability of wireless remote querymonitoring;

FIG. 11D depicts an artistic illustration of the recognition and bindingof pathogenic bacterial cells by robust and naturally occurringantimicrobial peptide based biorecognition moieties assembled on thegraphene nanomaterial transducer;

FIGS. 12A-F depict material integration and characterization for thebiosensor;

FIG. 12A depicts an image of a large area graphene nanotransducertransfer printed on to bioresorbable platform of silk fibroin

FIG. 12B depicts an image of the passive wireless telemetry systemconsisting of a planar meander line inductor and interdigitatedcapacitive electrodes integrated on to the graphene/silk nanocomposite;

FIG. 12C depicts an image of a large area graphene transducer/goldelectrodes bioresorbed on to the surface of human molar;

FIG. 12D depicts an image of a graphene based wireless elementbioresorbed on to tissue;

FIG. 12E depicts a graph of the raman spectra of surface grapheneintegration;

FIG. 12F depicts a graph of the raman spectra of the graphene integratedon to the tooth surface through silk bioresorption;

FIGS. 13A-C collectively illustrate single bacterium detection using thebiosensor;

FIG. 13A depicts, in the upper section, a graph of resistance datarecorded during the binding and unbinding of single E. coli O:157 H7bacterial cell over time on the graphene biosensor; the lower sectionshows simultaneous optical florescent photographs of the event;

FIG. 13B shows a schematic of a synthetized peptide consisting of abiocombinatorially derived graphene binding motif coupled to naturallyoccurring antimicrobial peptide (O-HP) based robust moieties atriglycine linker (-GGG-);

FIG. 13C depicts, in the upper section, a graph of resistance datarecorded during the binding of single E. coli O:157 H7 bacterial cellover time on the graphene biosensor; the lower section showssimultaneous optical florescent photographs of the event;

FIGS. 14A-F depict the wireless remote monitoring of S. aureus and H.pylori;

FIG. 14A shows the optical image of graphene based wireless sensorintegrated on to the surface of an I.V. bag;

FIG. 14B depicts a bar graph of the bandwidth the sensor resonancecorresponding to various concentrations of the S. aureus cells incubatedon the biosensor surface and dried;

FIG. 14C depicts a line graph of the percentage change in grapheneresistance versus log of concentration of S. aureus;

FIG. 14D depicts an optical image of the graphene based wireless sensorbioresorbed on to the surface of a bovine tooth;

FIG. 14E depicts a line graph of the percentage change in grapheneresistance versus time up on the exposure of a sample of 100 cells perμL H. pylori in DI water; and

FIG. 14F depicts a line graph of the percentage change in grapheneresistance versus log of concentration of H. pylori; error bars showstandard deviation (N=3).

DETAILED DESCRIPTION OF THE INVENTION

Definitions

As used herein, the phrase “low frequency” refers to a frequency below 1MHz.

As used herein, the phrase “high frequency” refers to a frequency above20 MHz.

As used herein, the phrase “electrical conductivity” refers to theelectrical resistance and/or impedance of a material.

As used herein, the initials “PCR” refers to polymerase chain reaction.

As used herein, the acronym “ELISA” refers to enzyme-linked immunoassay.

As used herein, the initials “AMP” refers to antimicrobial peptides.

As used herein, the acronym “NEMS” refers to nanomechanical cantileversensing.

As used herein, the acronym “SERS” refers to surface-enhanced Ramanspectroscopy.

As used herein, the initials “LPS” refers to lipopolysaccharide.

As used herein, the acronym “SNAP” refers to superlattice nanowirepattern transfer.

As used herein, the initials “NW” refer to nanowires. Similarly, as usedherein, the initials “SiNW” refer to silicon nanowires.

As used herein, the initials “FMOC” refer to fluorenylmethoxycarbonyl.

As used herein, the initials “O-HP” refer to Odorranin-HP, anantimicrobial peptide isolated from the skin of the frog speciesOdorrana grahami.

As used herein, the initials “IMA” refer to interdigitatedmicroelectrode array.

As used herein, the initials “PECVD” refer to plasma enhanced chemicalvapor deposition.

Current methods for detecting pathogenic bacteria include enzyme-linkedimmunoassay (ELISA), and polymerase chain reaction (PCR). In the formercase, the assays exploit antibodies as molecular recognition elementsdue to their highly specific targeting of antigenic sites. However,antibodies lack the stability needed to detect pathogenic species underharsh environments, reducing the shelf life of antibody functionalizedsensors. The high specificity of antibody-antigen interactions alsorequires a one-to-one pairing of antibody-based sensors for each targetto be detected. Nucleic acid probe-based techniques such as PCR canreach single-cell detection limits, yet require the extraction ofnucleic acids and are limited in portability.

By contrast, the ease of synthesis and intrinsic stability ofantimicrobial peptides (AMPs) render them particularly good subjectmatters for use as molecular recognition elements in bioelectronicsensing platforms. Anti-microbial peptides are a class of biomolecules,which appear in multiple niches in nature including the skin of higherorganisms and the extracellular milieu of bacteria as the primary lineof defense against bacteria and fungi. It is of particular note thatAMPs are much more stable than typical globular proteins explaining whythey can be continually exposed to the environment and are exceptionallyefficient at fending off bacterial infection. Indeed, some cationicantimicrobial peptides have shown activity toward pathogenic bacteriaunder harsh environmental conditions such as thermal(boiling/autoclaving) and chemical denaturants. The replacement ofcurrent antibody based affinity probes with more stable and durable AMPsin biological sensors may thus help to increase the shelf-life ofcurrent diagnostic platforms. A final major advantage of AMPs asrecognition elements stems from their semi-selective binding nature tothe target cells, affording each peptide the ability to bind to multiplepathogenic cells.

The durability and ruggedness of AMPs has led others to utilize AMPs asan enhancement to various devices to provide antimicrobial features.U.S. Patent Application Publication Nos. 2005/0065072 to Keeler et al.,2004/0126409 to Wilcox et al., and 2007/254006 to Loose et al. andEuropean patent No. EP 0 990 924 to Wilcox et at (all of which arehereby incorporated by reference in their entireties) describe variousmethods of attaching antimicrobial peptides to a variety of substratesfor the purpose of producing a device with antimicrobial properties.

A number of methods have been successful at detecting bacteria includingnanomechanical cantilever sensing (NEMS), surface-enhanced Ramanspectroscopy (SERS), and quartz crystal microbalance based sensors.Similarly, recent attempts have utilized AMPs as biorecognition elementsin fluorescent-based microbial detection with achievable detectionlimits of 5×10⁴ cells/mL. There have been previous attempts to create abiosensor that utilizes AMPs to detect bacteria. Notably, Kulagina etal. report the use of AMPs in a biosensor in U.S. Publication2006/0281074. However, it is worth particularly pointing out thatKulagina et al. used a florescent-based detection method which lacks theportability and near instantaneous results that are desirable in abiosensor. The biosensor of Kulagina et al. also does not permit realtime testing of sample in a flow-through system, another desirablefeature.

Detectors with real-time capabilities have been shown by others.However, these detectors lack the sensitivity and versatility of thebiosensor described herein. For instance, low field, low frequency,dielectric spectroscopy, as will be described further herein, haspreviously been used to analyze biologic material, see Prodan et al.(2004) J. Applied Physics 95(7): 3754-3756, hereby incorporated byreference in its entirety. However, Prodan et al. lack sensitivity for aparticular target. Prodan et al. do not use a biologic binding moleculeto detect cells merely the difference in impedance of differentfrequencies of current through a cell suspension to determine cellconcentration. It is also worth noting that the power source used in thestudy by Prodan et al. is hard wired into the detector. The detector ofProdan et al. lacks the specificity, portability and variability thatwould be desirable in a biosensor.

The development of an “all-in-one” solution which combines a high degreeof portability, robustness, sensitivity, and selectivity towardpathogenic strains remains challenging. There exists a need for a highlyportable biosensor device of pathogenic strains that is robust,sensitive and selective. It is also highly desirable if the biosensorcould be configured to operate on a variety of power sources, includingDC power, AC power and battery-free power. It further desirable if thebiosensor was scalable, so that it could function at laboratory size,handheld size and at nanometer size.

Biosensor

A label-free electronic biosensor based on the hybridization of theantimicrobial peptide with an electrode array for the sensitive andselective detection of antimicrobial peptide targets via impedancespectroscopy is disclosed. Specifically demonstrated is a label-freeelectronic biosensor based on the hybridization of the antimicrobialpeptide magainin I with interdigitated microelectrode arrays for thesensitive and selective detection pathogenic bacteria via impedancespectroscopy. Furthermore, it is contemplated that the combination ofcompact, naturally bioselective AMPs with microcapacitive sensorsrepresents a highly robust and portable platform for fundamentaldevelopment of AMP-bacteria interactions, and for portable infectiousdisease threat agent signaling. It is demonstrated that a bioresorbablesensor for highly sensitive and selective detection of biologicalanalytes is possible through the synergistic integration of the smartproperties of selectivity in biological recognition possessed bynaturally occurring antimicrobial peptides with the high sensitivity oftwo dimensional graphene nanomaterial transducers on a biocompatiblesilk fibroin substrate.

The biosensor's sensitivity is especially suited to detection ofinfectious agents. One aspect that makes the biosensor especially suitedto detect infectious agents is its capability to function in virtuallyany location. It is demonstrated that some embodiments of the biosensorare suitable for field use, for instance in a reservoir, whereinfectious agents might enter a population's water supply. It is furtherdemonstrated that some embodiments are suitable for use at the point ofinfection in vivo. It is demonstrated that certain embodiments of thebiosensor can be implanted on and used on the surface of a tooth. It iscontemplated that the biosensor can be used on nearly any biologicalsurface it can be brought in contact with including, but not limited totooth, bone, skin, tissue, hair, nail, cornea, gum, tongue, palate,brain, heart, lung, membrane, leaf, root, bark, fur, feather, chiton andscale. It is further contemplated the various embodiments of thebiosensor allow the biosensor to be used in nearly any environment,including but not limited to: in soil, in water, airborne, and in vivo.For instance, it is contemplated that nanoscale embodiments of thebiosensor could be injected subcutaneously or intravenously for useinside a human or other species. It is further contemplated that thebiosensor could be embedded on a substrate and swallowed like a pill asanother use in vivo.

Another aspect that makes the biosensor especially suited to detectinfectious agents is its high sensitivity. It is demonstrated that thebiosensor matches the sensitivity of PCR, in that it can detect levelsas low as a single infectious unit (see Example 10). Yet, unlike currentor imagined applications of PCR, the biosensor can provide this highdegree of sensitivity in real time and, as stated above, any location.This ability allows the biosensor to be an effective tool against highlyvirulent species or strains.

Every pathogen and strain of pathogen can be described in terms of itsminimum infectious dose (MID) to cause disease in a particular species,or subclass within a species. Some pathogens have a large MID for aparticular species, i.e., a dose of a large number of those pathogens isrequired for the clinical observation of disease. For instance, inRhesus monkeys as many as 10⁵ infectious units of H. Pylori are requiredfor the onset of disease (Solnick et al. (2001) Determination of theInfectious Dose of Helicobacter pylori during Primary and SecondaryInfection in Rhesus Monkeys (Macaca mulatta) Infection and Immunity, 69(11): 6887-6892 DOI: 10.1128/IAI.69.11.6887-6892.2001, herebyincorporated by reference in its entirety). For other infectious agentsin other species the number of units required to cause disease is muchless. Highly virulent infectious agents have a small MID against themost susceptible of hosts. For instance, as little as 10 viable cells ofShigella flexneri, Shigella sonnei or Shigella dysenteriae are needed tocause bacillary dysentery in humans (Schaad U. B. (1983) Which Number ofInfecting Bacteria is of Clinical Relevance? Infection 11, Suppl. 2,S87-S89, hereby incorporated by reference in its entirety). Therefore,it is especially necessary for biosensor to be able to detect as littleas 10 or fewer infectious agents to be effective as an alert to the mostvirulent of agents.

Transduction of Binding Event

Among the various label-free signal transduction platforms that areknown, impedance spectroscopy is most ideal due to its simpleinstrumentation, ease of device assembly, and adaptability tomultiplexed lab-on-a-chip applications. A microcapacitive sensor detectsimpedance changes in the dielectric properties of an electrode surfaceupon analyte binding, where the variation in the impedance is directlyproportional to the activity of analyte binding.

FIGS. 1A-1C show schematics of a measurement setup. In this example, anAC voltage applied to the electrodes produces both conduction anddisplacement current through the sample. The real and imaginary parts ofthe transfer function V2(w)/V1(w) are proportional to the conductivityand the dielectric constant, respectively. The output of the signalanalyzer is applied to one of the capacitive electrodes through R1. Theother electrode of the capacitive sensor is connected to the negativeinput of the amplifier A2, which holds the electrode at groundpotential. As a result, the current I that flows through the sensorproduces a voltage V2 which is equivalent to the product of I and thesample impedance Z. The value of voltage drop V1 is equal to the productof I and R2. Therefore, the transfer function of the system is given by:

$\frac{V_{2}(\omega)}{V_{1}(\omega)} = \frac{Z}{R_{2}}$

Where Z is the overall impedance of the sensing system. The purpose ofR1 is to provide an upper limit for the current I as the impedance Zbecomes smaller at higher frequencies. The unity gain amplifier A1provides buffering so that the input impedance of Channel 2 does notaffect the voltage drop across the sample.

A high density interdigitated microelectrode array can be used for thedetection of bacterial cells. The exposure of the magaininfunctionalized sensor to bacterial cells results in the binding on thecells on the electrode surface. Bacterial binding of bacterial cellscauses change in the impedance measured across the electrodes. FIG. 1Ashows an equivalent circuit of the microelectrode solution interfacebefore the binding bacterial cells. A display unit 14 is coupled to thebiosensor anode and cathode. The display unit includes a readout 16configured to indicate the conductivity, e.g., resistance and/orimpedance, of the biosensor. It should be understood that circuitsdisclosed herein may also include a power source as generally shown byblock 15. C_(DL) represents the capacitance due to the electrical doublelayer between the electrode and the buffer solution, C_(Di) representsthe dielectric capacitance, and R_(Buffer) the bulk resistance of thebuffer solution. A parasitic capacitance from the oxide layer betweensilicon and gold is shown as C_(PAR). FIG. 1B shows a simplified circuitdiagram of the system after bacterial binding to the AMP functionalizedelectrodes. The modification in the interface impedance due to thebacterial impedance consists mainly of the capacitance of the cellmembrane C_(CM), the resistance of the membrane R_(CM), and theresistance of the cytoplasm R_(Cyt), as shown. The represented model hastwo parallel branches, a dielectric capacitance branch and an impedancebranch. At high frequencies, the total impedance of the system Z will bedominated by the dielectric capacitance of the medium, and thecontributions from the electrical double layer capacitance and the bulkmedium resistance will be minimized. At lower frequencies (<1 MHz),current does not flow through the dielectric capacitance branch, and thebacterial cells bound to the electrodes add different impedancecomponents in series to the impedance branch.

Sensitivity Measurements

As a test case for the demonstration of an AMP-based label-free,electronic biosensor, the targeting of microbial cells by magainin I wasconducted using impedance spectroscopy. FIGS. 2A-2C schematicallyoutline the sensing platform. AMPs are first immobilized onmicrofabricated interdigitated gold electrodes, anode 10 and cathode 12(FIG. 2A). Magainin I was acquired with an additional cysteine residueat the C-terminus (FIG. 2B), allowing for facile and site specificcovalent attachment to the gold electrodes. Next, heat-killed bacterialcells were injected and incubated with the AMP-modified electrodes. Ifthe bacteria are recognized by the AMPs, binding will occur (FIG. 2C),leading to dielectric property changes which can be monitored via aspectrum analyzer. The impedance was measured over a frequency range of10 Hz to 100 kHz. FIG. 2D shows an optical micrograph of the device,which is made using standard microfabrication techniques.

Sensitivity of microbial detection is a key determinant for utility ofthe biosensor. To this end, the sensitivity of themagainin-functionalized microelectrode array in detecting bacterialcells was determined using impedance spectroscopy. FIGS. 3A-3B show theresults of measurements performed after incubation of the immobilizedAMPs with pathogenic E. coli O157:H7 cells in concentrations rangingfrom 10³ to 10⁷ CFU/mL. A “blank” device with no immobilized AMPs wasalso tested for comparison. FIG. 3A shows that at low frequencies, thedifferent concentrations of bacterial cells have the effect ofincreasing the impedance in proportion to the number of cells present inthe sample for concentrations greater than 10² CFU/mL. As the frequencyincreases, the contribution to the impedance from the bacterial cellsdecreases, leaving only the dielectric relaxation of small dipolesincluding water molecules in the buffer solution to affect the measuredimpedance. FIG. 3B depicts the impedance change at the fixed frequencyof 10 Hz. The variation in the impedance is directly proportional to thenumber of bacterial cells bound to the immobilized AMPs and manifestedin a logarithmic increase with respect to serially diluted bacterialconcentrations. Significantly, the detection limit of response of thehybrid AMP-microelectrode device to E. coli is found to be 10³ CFU/mL (1bacteria/μL), as the dielectric polarization of the bacterial cellsbound to the surface of the electrodes begins to have an effect on thebase impedance of the capacitive electrodes at this concentration (seeFIG. 4). This sensitivity limit is clinically relevant and comparesfavorably to AMP based fluorescent assays [5×10⁴ CFU/mL] and toantibody-based impedance sensors. This lowest limit of detection appearsto be limited by the presence of impedance due to the electrical doublelayer resulting from the electrode polarization effect at lowfrequencies.

To gain further insight into the sensitivity of the magainin I AMPtoward E. coli, AMPs were immobilized “upside-down” via incorporation ofa cysteine residue at the N-terminus. The binding affinities of magaininI immobilized via cysteine residues at the C-terminus and N-terminuswere compared and co-plotted in FIGS. 3A and 3B. Considerably reducedbinding activity was observed for magainin immobilized via theN-terminus compared to C -terminal immobilization.

Without being bound by speculation, it is thought this reduction in thebinding affinity is likely due to the diminished exposure of the targetbacteria to the amine-containing residues near the N-terminus. Thisobservation supports the hypothesis that the initial interaction ofα-helical AMPs with the membranes of the target bacteria occurs viaelectrostatic attraction of positively charged amino acids on the AMPwith negatively charged phospholipids in the bacterial membrane. Indeed,it has been previously shown that amino acid omissions in the N-terminalregion of magainin result in the complete loss of antimicrobialactivity, whereas analogs with omissions in the C-terminal regionexhibited equal or increased activity.

Finally, the effect of varying the surface density of the immobilizedAMPs on the detection of bacterial cells was investigated (see FIG. 5).The response of the biosensor towards target cells was found to increasemonotonically with increasing concentration of immobilized magainin.

In an effort to determine the optimal performance of the biosensor, thevalue of varying the surface density of the immobilized AMPs on thedetection of bacterial cells was explored. Different concentrations ofC-terminal cysteine labeled magainin I were immobilized on the electrodesurface. The impedance response resulting from binding of pathogenic E.coli O157:H7 cells (10⁷ CFU/mL) to different densities of immobilizedAMPs were recorded. The response of the sensor at 10 Hz is plotted inFIG. 5. The immobilized peptide film was also analyzed via fluorescentmicroscopy by labeling the peptides with fluorescein isothiocyanate(FITC). The ability to capture the target bacteria was found to bestrongly dependent on the immobilization density of the magainin on thesensor surface. This supports the hypothesis that the initialinteraction between the cationic AMPs and the target species occursthrough electrostatic interaction. This also suggests that theminimization of diffusion and steric hindrance, which usually affect thebinding kinetics, do not play a significant role in the immobilizedAMP-bacteria interactions.

Selectivity Measurements

As a next step, the selectivity of the AMP-functionalized biosensorstoward various bacterial species was determined. Specifically, thebinding behavior of AMPs was probed toward: 1) Gram-negative pathogenicE. coli O157:H7, 2) the non-pathogenic E. coli strain ATCC 35218, 3)Gram-negative pathogenic Salmonella typhimurium, and Listeriamonocytogenes, a Gram-positive pathogen. Collectively, these studieselucidate the matrix of selectivity as it depends on Gram-negative vs.Gram-positive species, and pathogenic vs. non-pathogenic strains. Theselectivity was first investigated using fluorescent microscopy methods,by staining bacterial cells and optically mapping their binding densityto gold films hybridized with AMPs. FIGS. 6A-D shows the discriminativebinding pattern of immobilized magainin I to various bacterial cells(all 10⁷ CFU/mL) stained with propidium iodide (PI) nucleic acid stain,as well as the surface density of the bound bacterial cells. Likewise,FIG. 7A plots the electrical response of the AMP-biosensor against thesevarious species as a function of the interrogating frequency, and FIG.7B plots the response at 10 Hz.

As a proof of principle, inspection of the fluorescent images andsurface density plots agree qualitatively with the response of the AMPelectrical biosensor and reveal the following performancecharacteristics. First, magainin I exhibits clear preferential bindingtoward the pathogenic, Gram-negative species E. coli and Salmonella,relative to the Gram-positive species Listeria, with a two order ofmagnitude impedance difference at 10 Hz (FIG. 7B). This selectivity wasparticularly enhanced for pathogenic E. coli, which showed slightlylarger response relative to Salmonella. Next, inter-bacteria straindifferentiation between pathogenic and non-pathogenic bacteria isdemonstrated by the ability of the sensor to selectively detectpathogenic E. coli relative to the non-pathogenic strain, again with anearly two order of magnitude impedance difference at 10 Hz. Finally,the response of the sensor to all microbial species was larger than theresponse of the “blank” biosensor which was not functionalized with AMP.

The observed specificity differences provide support for a balancebetween electrostatic and hydrophobic interactions that is theunderlying hypothesis of the mechanism of binding to bacterial cells byAMPs. In the case of magainin I, the difference in the membranestructures of Gram-negative vs. Gram-positive bacteria accounts for thedifferential selectivity. Gram-negative bacteria possess an outermembrane with negatively charged lipopolysaccharide (LPS)—the first siteof encounter for AMPs—and a thin peptidoglycan layer. In contrast,Gram-positive bacteria lack the LPS containing outer membrane andinstead possess a thick peptidoglycan layer and teichoic acids. Further,although both pathogenic and non-pathogenic E. coli cell walls containLPS, the LPS of pathogenic strain includes O-antigens, which arehydrophilic branched sugar side chains. These O-antigens form theoutermost portion of the polysaccharide chain and are thought to enhanceelectrostatic and hydrogen bonding. This ability of magainin I toselectively prefer Gram-negative species, and pathogenic versusnon-pathogenic strains of E. coli, agrees with several other bacteriaadhesion studies. Therefore, it is expected that AMPs will perform aspart of the biosensor as the AMPs are observed to function in vivo.

Real-Time Detection

To simulate the use of the AMP microelectrodes in everyday applications,such as direct water sampling, the biosensor response tested to performin real time, as shown in FIGS. 8A-8D. First, a microfluidic cell wasbonded to the interdigitated biosensor chip (FIG. 8A), such that theelectrodes were perpendicular to the direction of the sample flow (FIG.8B). Next, fluid was injected using a syringe pump connected to theinlet port, and allowed to flow through to the outlet port, at a flowrate of 100 μL/min. The flow cell was first flushed with buffer toestablish a baseline. Various dilutions (10⁴-10⁷ CFU/mL) of pathogenicE. coli cells in PBS were then injected to the channel at a reduced flowrate of 5 μL/min for 30 min. For example, FIG. 8C shows themicroelectrode array after exposure to 10⁷ CFU/mL bacterial cells.Simultaneously, the impedance response was continuously monitored duringthe sample flow-through process (FIG. 8D). Significantly, all samplesproduced a measurable response relative to the control sample within 5minutes, with the highest concentration sample yielding a responsewithin 30 seconds; these responses saturated after ca. 20 min. Althoughthese results bode well for the implementation of this sensor incontinuous monitoring of flowing water supplies, it should be noted thatfor the same concentration of bacterial cells, the response of thesensor under flow-through conditions was found to be comparatively lowerthan the response after static incubation. This difference betweendetection in flowing and static water sample has also been reported influorescent based assays. This difference is attributed to the opposingeffects of shear and mixing on the binding kinetics.

In summary, coupling of AMPs with microcapacitive biosensors hasresulted in the development of a portable, label-free sensing platformfor the detection of infectious agents. The achievable sensitivityapproached 1 bacterium/μL—a clinically relevant limit—and the AMPsallowed for sufficient selectivity to distinguish pathogenic andGram-negative bacteria, while retaining broadband detectioncapabilities. Finally, real-time sensing results demonstrated thecapability of the relatively simple impedance-based transductionarchitecture to directly detect bacteria, demonstrating an improvementto traditional antibody based immunoassays. These results provide asignificant positive improvement on the use of pathogenic sensors totest and monitor bacteria in reservoir water, or for use as biologicalthreat agent detection systems. It is contemplated that this biosensorcan be utilized for the detection of bacteria in real water samples.

It is further contemplated that these biosensor configured with AMPpeptides coupled to silicon nanowire sensors will produce sensor withsignificantly enhanced sensitivity and a size of the nanometer scale.

Antimicrobial Peptides

Small, cationic antimicrobial peptides (AMPs) are naturally occurringantibiotics of the innate immune system. AMPs are widely distributed inanimals and plants and are among the most ancient host defense factors.Their spectrum of activity may include Gram-positive and Gram-negativebacteria as well as fungi cancer cells, certain viruses and somemulticellular animal cells.

The bioactivity of AMPs toward microbial cells is classified into groupsaccording to their secondary structures. Many AMPs adopt amphipathicconformations that spatially cluster hydrophobic from cationic aminoacids, thereby targeting the negatively charged head groups of lipids inthe bacterial membrane. In contrast, the membranes of plants and animalsseclude negative charges to the inner leaflet, and contain cholesterolswhich reduce AMP activity. By aiming at the very foundation of thebacterial cell membrane, and remaining generically unrecognizable toproteases, AMPs as antibiotics have remained remarkably free of acquiredresistance. It is therefore anticipated that AMPs represent a class ofbiomolecules that will continue to bind to a target cell membranesdespite genetic drift and mutation within the population. AMPs can alsobe expected to bind to targets cell membranes across a wide range ofgeographic areas wherein there many be variation local of strains andspecies of target cell membranes.

Among AMPs, linear cationic peptides such as magainins are particularlysuited for use with the biosensor because of their small molecular sizeand intrinsic stability. In particular, the positively charged AMPmagainin I (GIGKFLHSAGKFGKAFVGEIMKS) (SEQ ID NO: 1) binds mostselectively to the bacterial cell E. coli O157:H7 as a precursor tobactericidal activity. Magainin I also displays broad spectrum activitytoward other Gram-negative bacteria, which comprise the majority ofpathogenic infection in humans. Therefore, Magainin I is an ideal testcase for demonstrating the capabilities of the biosensor.

However, the biosensor is not limited for use with Magainin I. Virtuallyany AMP, modified AMP, or AMP binding motif, can be used with thebiosensor. The criteria for suitability with the biosensor is (1) aknown target to which the candidate AMP selectively binds and (2) theability to immobilize the AMP, modified AMP, or AMP binding motif, to anelectrode without disabling the ability of the peptide to selectivelybind its target.

There exists an AMP database with which it is possible to search forknown AMPs with desirable characteristics. Currently, the AMP databaselists 18 AMPs with known affinity to E. coli. These peptides are shownin Table 1. The database can currently be found that the followinginternet address: http://aps.unmc.edu/AP/main.php

TABLE 1 BD SEQ REF ID Name Sequence ID NO AP00580 Nigrocin-2GRbGLFGKILGVGKKVL  2 CGLSGMC AP00678 K9CATH RLKELITTGGQKIG  3EKIRRIGQRIKDFF KNLQPREEKS AP00807 Enterocin E-760 NRWYCNSAAGGVGG  4AAGCVLAGYVGEAK ENIAGEVRKGWGMA GGFTHNKACKSFPG SGWASG AP00846 Mundticin KSKYYGNGVSCNKKGC  5 SVDWGKAIGIIGNN SAANLATGGAAGWK S AP00964 Dermaseptin-L1GLWSKIKEAAKAAG  6 KAALNAVTGLVNQG DQPS AP01373 Human TC-1 AELRCMCIKTTSGI 7 HPKNIQSLEVIGKG THCNQVEVIATLKD GRKICLDPDAPRIK KIVQKKLAGDES AP01374Human TC-2 NLAKGKEESLDSDL  8 YAELRCMCIKTTSG IHPKNIQSLEVIGKGTHCNQVEVIATLK DGRKICLDPDAPRI KKIVQKKLAGDES AP01380 TBD-1 YDLSKNCRLRGGIC 9 YIGKCPRRFFRSGS CSRGNVCCLRFG AP01382 2B5B, TEWP EKKCPGRCTLKCGK 10HERPTLPYNCGKYI CCVPVKVK AP01402 Ocellatin-V1 GVVDILKGAGKDLL 11AHALSKLSEKV AP01403 Ocellatin-V2 GVLDILKGAGKDLL 12 AHALSKISEKV AP01404Ocellatin-V3 GVLDILTGAGKDLL 13 AHALSKLSEKV AP01405 LeptoglycinGLLGGLLGPLLGGG 14 GGGGGGLL AP01407 SgI-29 HNKQEGRDHDKSKG 15HFHRVVIHHKGGKA H AP01578 Myxinidin GIHDILKYGKPS 16 AP01591 cBD-1KCWNLRGSCREKCI 17 KNEKLYIFCTSGKL CCLKPKFQPNMLQR AP01648 PelteobagrinGKLNLFLSRLEILK 18 LFVGAL AP01753 Vejovine GIWSSIKNLASKAW 19NSDIGQSLRNKAAG AINKFVADKIGVTP SQAASMTLDEIVDA MYYD

It is contemplated that any antimicrobial peptide can be used withcertain embodiments of the biosensor. It is further contemplated thatcertain embodiments of the invention could use modified antimicrobialpeptides, or chimeric peptides containing antimicrobial peptide bindingmotifs and, combinations thereof.

Fields of Use

Biosensors utilizing electrically coupled antimicrobial peptides can beuse any field where detecting a target of a known antimicrobial peptideis desirable. These biosensors can be configured with any knownantimicrobial peptide, modified AMP, AMP binding motif or a combinationthereof. Therefore, they can be used as a diagnostic or monitoring toolin any situation where the detection and analyzes of an antimicrobialpeptide target is desirable. The possible uses of the biosensor includebut are not limited to applications such as diagnosis, food and waterquality monitoring and microbial contamination monitoring in hospitals.Non-limiting examples include use of the biosensor in a hospitallaboratory to screen a patient's bodily fluid for infection by apathogen; use in the field to test a body of water for the concentrationof a class or specific bacteria; and interfacing the sensor withpatient's tooth through bioresorption of the temporary substrate toenable wireless monitoring and detection of specific pathogenicmicroorganisms or disease markers in saliva or breath.

Size

As shown in the above examples of use biosensors utilizing electricallycoupled antimicrobial peptides can be tailored to a size for almost anyproposed use. For instance, if the end use is in a laboratory it may bedesirable to use a biosensor of bench-top size. This embodiment of thebiosensor could have an AC power source directly coupled to the circuitcoupled to immobilized AMPs. Also, in this embodiment of the sensor themicrocapacitive sensor may also be directly coupled to the same circuitas the AMPs. It is also contemplated that some embodiments may have anarray of circuits each with particular a class or species of immobilizedAMPs electrically coupled to a particular circuit. In this embodiment ofthe biosensor, a plurality of analytes can be detected and theconcentrations thereof determined. These different circuits each withits particular AMP could be situated in different channels so that flowof a sample could be directed to the particular AMP circuit.Alternatively, multiple circuits could be situated on a single siliconchip, or similar substrate, with exposure of the sample to all thecircuits on the chip.

If the end use of the biosensor is in the field, it may be desirable tohave a handheld size device. Such an embodiment of the biosensor couldbe a portable device. It is contemplated that a battery, typically a DCpower source, could power a handheld embodiment of the biosensor.Similar to the AC powered unit described above, the handheld embodimentcould have the microcapacitive sensor electrically coupled to the samecircuit as the immobilized AMPs. The held-hand size would still permit adesign with multiple circuits of different classes or species of AMPs tobe combined in a single device.

If the end use of the biosensor is to be in vivo it is desirable thatbiosensor be sized on order of micrometers or preferably nanometers.

Recently, it has been shown that many types of organic devices includingtransistors, sensors, and photovoltaic cells can be fabricated on bothnatural and synthetic flexible polymers including, but not limited to,poly(ethylene terepthalate), poly(imide), poly(ether sulfone),cellulose, paper, silk, silk fibroin and combinations thereof. Thesepolymers are suitable to act as a substrate for immobilized AMPs.Furthermore, it is contemplated that the immobilized AMPs on thesesubstrates could be electrically coupled into a circuit using nanowires.These nanocircuits electrically coupled to immobilized AMPs could beplaced anywhere where it would be desirable to detect the analyte of theimmobilized AMP. As non-limiting examples, nano-biosensors could beplaced in catheters to test for pathogens, or directly on food to detectfor contamination and/or spoilage, or in an experimental animal subjectto detect bacteria of research study, or in human patients to detectcancerous or precancerous cells or any other threat to the human body,or in the human bloodstream to detect glucose concentration withoutbloodletting or on the blade of fans to detect air quality, oralternatively nanoscale embodiments could be airborne to detect or seekout air pollutants. The minute size of nanoscale biosensor combined withthe described wireless telemetry system allows an embodiment of thebiosensor to be used to detect virtually an analyte anywhere.

The substrates listed above in certain situations can be bioresorbed.Gradual bioresorption of the supporting substrates (for instance silkfibronectin) leaves the ultra-thin sensors intimately in place. Thenanometer size scale of the sensor allows it to conform to thecurvilinear surfaces biological tissues, teeth or any arbitrarysubstrate.

Highly sensitive and selective biorecognition can be achieved throughintegration of naturally occurring antimicrobial peptides basedbiorecognition moieties with highly sensitive graphene transducers.

Such embodiments could be operated with battery-free wirelessmechanisms. Battery free operation can be achieved through the magneticcoupling of a planar coil antenna interfaced with the nanomaterialtransducers to a remote reader, also enabling subsequent wirelessred-out.

Coupling of Peptides to Nanowires.

Silicon nanowires (SiNW) surfaces terminate in intrinsic silica, whichhas a well-established chemistry, that permits nanowire (NW) surfacemodification without strongly affecting the semiconducting core. TheSiNWs can be fabricated using the superlattice nanowire pattern transfer(SNAP) method, which can be harnessed to produce highly regular arraysof virtually any material that can be obtained as a high quality thinfilm. A typical NW array may comprise 400 high aspect ratio (>10⁵), 16nm wide SiNWs at a pitch of 33 nm, with a p-type doping level of˜10×18/cm³. These NWs perform as excellent field-effect transistors onboth solid and flexible plastic substrates, with mobilities comparableto that silicon. The SiNWs can be fashioned into sensor devices viaconventional microfabrication techniques. A contact metal layer of 1000Å Ti can be uniformly evaporated across the entire chip and subsequentlypatterned via photolithography and HF etching to form source/drainfinger electrodes across the SNAP wire array. The NW film can be thensectioned into individual sensor elements using photolithography andetching. As an example of sensitivity, SNAP SiNW sensors are capable ofdetecting parts per billion levels of an analyte. Peptides can besynthesized by the fluorenylmethoxycarbonyl (FMOC) solid phase peptidesynthesis method, in which FMOC-protected amino acid residues aresequentially linked on a solid bead support via repeated cycles ofcoupling and deprotection. The peptides remain immobilized on the beadsuntil cleaved by trifluoroacetic acid. Peptides can be subsequentlypurified to >95% by HPLC using a C 18 semipreparative column. Peptidescan then be immobilized onto the NWs using amide coupling. First, thenanowire surfaces are chemically modified by immersion of the chip in anamino silane (APTES) modifying reagent. Next, oligopeptides aresynthesized with the desired recognition sequences, plus an asparticacid “linking residue” tail at the carboxy terminus. The peptides can bedissolved in DMF, mixed with coupling reagents, and immediately injectedinto PDMS microfluidic chambers aligned to the device islands (FIGS.9A-9B). Once this coupling reaction is complete (2 h), the microfluidicchannels are removed and the chip is thoroughly rinsed to remove anyuncoupled peptide. Finally, the chip can be treated to a piperidinesolution to cleave the FMOC protecting group. X-ray photoelectronspectroscopy (XPS) measurements on silicon-on-insulator (SOI)\waferpieces, treated to identical surface reaction protocols as the SiNWsensors, can be employed to monitor this coupling chemistry.

SiNW Sensor Fabrication.

A typical set of NW sensors that can be employed for this biosensor isshown in FIGS. 10A-10F. The chip containing the SNAP wire arrays can betreated to mild O₂ plasma (300 mTorr, 30 W, 60 s), then immersed inbuffered oxide etch (BOE) for 3 s to remove oxides and promote theformation of ohmic contacts. Source and drain electrodes are formed byelectron-beam evaporating 1000 Å Ti uniformly across the chip, and thenpatterning the Ti through a photoresist mask (Shipley 1813, MicroChemCorp., Newton, Mass.) via wet etching (1:1:10 HF/H₂O₂/DI v/v, 5 s). Theresulting device channels are about 5 μm in length. A new photoresistmask can be applied to expose unwanted regions of the NW array forsectioning into device islands. The Si can be removed via reactive-ionetching (SF 6, 20 sccm, 20 mTorr, 30 W, 1 min), and the photoresist canbe removed in acetone.

Fabrication of P-Type Doped Snap Nanowires

Silicon nanowires can be fabricated from an intrinsic, 320 Å thick SOIfilm ({100} orientation) with a 2,500 Å buried oxide. The substrate canbe thoroughly rinsed and cleaned with deionized water then coated withp-type spin-on dopants. Dopants can be diffused into the SOI film usingrapid thermal processing at 800° C. for 3 min. Analyzes with four-pointresistivity measurements, correlated with tabulated values, shows atypical doping level of about 10¹⁸ cm⁻³. Separately, a superlatticeconsisting of 800 layers of alternating GaAs and AlxGa(1−x) As thinfilms can be prepared. The superlattice can be cleaved along a singlecrystallographic plane and thoroughly cleaned by sonicating in methanoland gentle swabbing. The exposed edge can be immersed in NH₃/H₂O₂/H₂O(1:20:750 v/v) for 10 s to selectively etch the GaAs regions (etch depthof about 30 nm). The resulting edge of the superlattice thereforeconsists of AlxGa(1−x) As plateaux separated by GaAs valleys. Pt metalcan be deposited using electron-beam evaporation onto the edge of theAlxGa(1−x) As ridges, with the edge of the superlattice held at a 45°angle to the incident flux of Pt atoms. The Pt-coated superlattice edgecan then be brought into contact with the doped SOI substratespin-coated (6,000 r.p.m., 30 s) with a thin-film PMMA/epoxy (1:50wt/wt). The superlattice/epoxy/SOI sandwich can be dried on a hot plate(150° C., 40 min), and the superlattice can be released by a selectiveetch in H₃PO₄/H₂O₂/H₂O (5:1:50 v/v, 4.5 h) solution, leaving a highlyaligned array of 400 Pt NWs on the surface of the SOI substrate. ThesePt NWs serve as protective masks for a reactive ion etch process toproduce aligned, single-crystal Si NWs (CF₄/He, 20/30 sccm., 5 mtorr, 40W, 3.5 min). The Pt NWs can be dissolved in aqua regia (30 min) toproduce an array of 400 Si NWs. Finally, the substrate can be cleaned inALEG-355 solution to remove residual epoxy.

Among the various label-free signal transduction platforms that havebeen investigated, impedance spectroscopy is promising due to its simpleinstrumentation, ease of device assembly, and adaptability tomultiplexed lab-on-a-chip applications. A microcapacitive sensor detectsimpedance changes in the dielectric properties of an electrode surfaceupon analyte binding, where the variation in the impedance is directlyproportional to the activity of analyte binding.

Graphene/Silk Based Bioresorbable Passive Wireless Sensor

The operation and some functionalities possessed by our sensingtechnology is schematically illustrated in FIGS. 11A-D. FIG. 11A depictsan artistic illustration of the graphene based highly sensitive passivewireless biosensor element on a bioresorbable substrate of silk fibroin.The biosensor includes a wireless telemetry device 18 coupled to asensor portion 20. In this example, the wireless telemetry device is aresonant circuit with an inductive coil. In this example the sensorportion 20 includes an immobilized peptide electrically coupled to ananode and a cathode. FIG. 11B illustrates the ability of theseultra-thin biosensors to be bioresorbed from the silk platform andintimately attached onto biological tissues, bones, or teeth through thedissolution of the supporting silk film for potential infectious agentmonitoring. The high surface area of the graphene and electrodes ensureshigh adhesive conformability to the curvilinear surfaces of biologicaltissues such as skin or bone. Specificity in biological recognition isachieved through the integration of designer bifunctional peptidesconsisting of robust and naturally occurring antimicrobial peptide basedbiorecognition moieties assembled on to the graphene surface throughcombinatorially derived graphene binding peptide. The modification ofantimicrobial peptides with peptide oligomers that bind to grapheneenables the facile and non-covalent functionalization of the graphenenanomaterial transducers without affecting its excellent electronicproperties. FIG. 11C illustrates the two other functionalities of thehybrid biosensor unit, particularly, battery free operation and remotewireless sensing capability. In this example, an antenna 22, e.g., aplanar coil antenna, is disposed in proximity to the wireless telemetrydevice 18. A display unit 24 is coupled to the antenna 22. The displayunit generally contains a transduction sensor configured to determinethe conductivity of the biosensor. Upon the recognition and binding ofspecific bacterial targets by the immobilized antimicrobial peptides(FIG. 11D), the electrical conductivity of the graphene film changes,which can be wirelessly monitored using the inductively coupled radiofrequency reader device 26. The presence of passive wireless componentsintegrated with the resistive transducer enables the wirelessinterrogation of the sensor and any change in the conductance of thegraphene layer is manifested as a modification of the fundamentalelectrical parameters of the oscillating circuit at resonance. Thisdevice can be used to remotely detect a wide variety of species,including, but not limited to, neuronal signaling networks, chemicalthreats such as TNT, bacteria on the skin or in the saliva viaantimicrobial peptides, and/or disease metabolites which appear in thebreath by implementing the sensor on a tooth.

Material Integration and Characterization

Some functionalities of the graphene/silk hybrid sensing elements arederived from the synergistic integration of the smart propertiespossessed by its component materials. The biosensor interfaces passivewireless graphene nanosensors with biological substrates viabioresorbable silk substrates, including, but not limited to bone,teeth, and tissue. Large area, functional graphene nanomaterialtransducers integrated with water soluble silk fibroin films of ca. 50microns thick through a simple transfer printing process (FIG. 12A). Itis possible for large-scale printing of graphene nanosensors onto silksubstrates. Such techniques are demonstrated in Mannoor, M. S., Clayton,J. D., Tao, H., Omenetto, F. G., McAlpine, M. C. (2011)Graphene/Silk-Based Bioresorbable, Passive Wireless Sensors, manuscriptsubmitted for publication, hereby incorporated by reference in itsentirety.

Electrode patterns are incorporated on to the silk-graphene compositefilms through a shadow mask assisted electron beam evaporation of gold.FIG. 12B shows image of a graphene/silk hybrid biosensor device. Thearchitecture consists of a parallel LRC resonant circuit with a goldinductive coil for wireless transmission, and interdigitated capacitiveelectrodes contacting sensitive graphene resistive sensors to form apassive wireless telemetry system, obviating the need for onboard powersources and any external connections. A full-wave electromagneticsimulation tool, Ansoft HFSS can be used to simulate and design variousgeometries of planar coil antenna and interdigitated capacitiveelectrodes.

The thin film sensing elements on silk platforms are then bioresorbedand intimately integrated on to a variety of substrates through thedissolution and removal of the supporting silk template. FIG. 12C showsthe image of graphene nanomaterial with pattered gold electrodesintegrated on to the surface of a human molar. Optical characterizationof the graphene transferred on to SiO₂ surface through the dissolutionof silk film revealed good structural integrity. FIG. 12D shows theimage of the graphene/silk hybrid sensor bioresorbed on a tissue.Complete dissolution of silk matrix in water leading to the attachmentof the graphene-Au electrode nanocomposite is observed with in a timeperiod of 15-20 minutes.

The electronic properties and structural information of the grapheneintegrated with the tooth surface is investigated using Ramanspectroscopy studies. FIG. 12E shows the Raman spectrum of a toothsurface before the integration of graphene nanosensor. FIG. 12F showsthe spectrum of the graphene nanosensor integrated on to the toothsurface via bioresorption of silk substrate.

Single Bacterium Detection

To demonstrate the response of the graphene-nanosensors towards singlebacterial cells and to characterize the sensitivity of the graphenetransducing element, simultaneous electrical and optical measurementswere conducted. Time dependent resistance data recorded simultaneouslywith optical measurements on graphene nanosensors showed discretechanges in the graphene resistance in the presence of a sample ofapproximately 100 bacterial cells per μL loaded on to the sensorelements by pipette (see FIG. 13A). Simultaneous collection ofelectrical and fluorescence data from the graphene sensors in thepresence of fluorescently labeled E. coli cells clearly indicate thatthe change in the resistance recorded corresponds to the binding andunbinding of the bacterial cell from the graphene surface. Theresistance of the graphene sensor remains at the baseline value untilthe bacterial cell diffuses to the sensor surface and drops down byapproximately 0.6 percentage of the original, once the bacterial cellbinds to the graphene surface.

The resistance of the sensor decreases from the baseline value up on thebinding of the negatively charged bacterial cell and returns back to theinitial value up on unbinding, indicating a p-type behavior of thegraphene transducer. The ability to specifically distinguish betweenvarious species of pathogenic bacteria is important in utilizing thehigh sensitivity of the graphene nanosensors in most medical andbiorecognition applications. In is contemplated that biocombinatorialscreening techniques such as phage display can be used to determinepeptide sequences which selectively bind to graphene. In order to impartselectivity to the sensing mechanism, graphene transducers werefunctionalized with antimicrobial peptide O-HP which has been identifiedto show specific activity toward both Gram-negative bacteria (E. coliand H. pylori) and Gram-positive bacteria (S. aureus). Facile andefficient assembling of the antimicrobial peptides on the graphenenanosensors were enabled by generating bifunctional peptides that linkthe graphene binding motif and O-HP separated by triglycine, resultingin the 38 amino acid sequence, HSSYWYAFNNKT-GGG-GLLRASSVWGRKYYVDLAGCAKA(SEQ ID NO: 20) (FIG. 13B). The triglycine linker provides flexibilityand independent accessibility of the bacterial cells towards the O-HP.In order to determine the effect of the immobilized AMPs on thebinding/unbinding properties simultaneous resistance and optical datawere recorded on graphene sensors functionalized with antimicrobialpeptides. Continuous binding of bacterial cell without unbinding for amuch longer time period compared to the duration of binding in the caseof non-functionalized graphene was observed (FIG. 13C). This shows thatthe immobilized peptide (O-HP) slows the unbinding kinetics of thebacterial cells. FIG. 13C inset shows the fluorescent image of thegraphene-interdigitated electrode nanosensor functionalized withFITC-labeled antimicrobial peptides.

For the purpose of determine the location of the binding motif withinthe peptide, a bifunctional peptide sequence with the antimicrobialpeptide in an inverted fashion was synthesized. Binding density analysiswith fluorescently labeled bacterial cells did not show any significantdifference in the binding affinity between the normal and invertedO-HPs, indicating that the sequences responsible for the bacterialbinding is present at the middle of the AMP sequence.

Wireless remote-query monitoring of S. aureus and H. pylori. Onefunctionality of the graphene/silk hybrid sensing element is thewireless remote-query capability. Staphylococcus aureus, a Gram-positivepathogenic bacteria found on skin flora and hospital environments, isnotoriously drug-resistant and responsible for over 500,000post-surgical wound infections in the US per year. To simulate the useof the graphene wireless sensor in hospital sanitation and biohazardmonitoring, the sensing elements were integrated on to an intravenous(I.V.) bag through the dissolution of the supporting silk substrate(FIG. 14A). S. aureus has been reported to survive up to 9 weeks onstandard plastic and similar dry hospital environments. To demonstratethe capability of the biosensor to detect S. aureus in typical hospitalconditions, 1 μL solutions containing various concentrations (10³-10⁸cfu/mL) bacterial cells were delivered to the biosensor and allowed themto dry on the biosensor surface for 30 min. The change in the grapheneresistance up on bacterial binding is wirelessly monitored as thebandwidth change in the sensor resonance curve. FIG. 14B plots thebandwidth of the sensor corresponding to the different concentration ofthe S. aureus cells incubated on the sensor surface. The percentagechange in graphene resistance is calculated from the bandwidth and isdepicted in FIG. 14C.

In order to demonstrate the performance of the sensor directlyintegrated with biological tissue, creating the possibility for on-bodyhealth quality monitoring, the wireless sensing elements werebioresorbed and integrated on to the surface of a tooth (FIG. 14D). Thisembodiment of the biosensor enables the remote monitoring of thepresence of infectious agents in saliva and disease metabolites inbreath. To this end, the response of the graphene hybrid sensors toHelicobacter pylori—a Gram-negative species found in the stomach andsaliva which is estimated to be responsible for the development of over90% of duodenal ulcers and stomach was analyzed. Real-time change ingraphene resistance up on exposure to various concentrations of H.pylori cells in DI water was monitored by recording the characteristicfrequencies at resonance. FIG. 14E depicts the real-time change ingraphene resistance up on the exposure to a 1 μL sample containing 100H. pylori cells. The sensor resistance is observed to be stabilizedafter around 10 to 15 min. The response of the graphene sensing elementto 1 μL of “blank” DI water without any bacterial cells is used as acontrol. FIG. 14F depicts the percentage change in resistance as afunction of bacterial concentration. A linear relationship is observedbetween the logarithm of bacteria concentration and the resistancechange up to a concentration of 10⁶ bacterial cells.

Summary of Nanoscale Biosensor

The biosensor in its nanoscale embodiment is the first example of apractical realization of the direct integration of functionalmaterial—graphene nanosensors—with human body to function as astandalone biological sensor. A bioresorbable sensor for highlysensitive and selective detection of biological analytes is realizedthrough the synergistic integration of the smart properties ofselectivity in biological recognition possessed by naturally occurringantimicrobial peptides with the high sensitivity of two dimensionalgraphene nanomaterial transducers on a biocompatible silk fibroinsubstrate. The incorporation of a parallel resonant circuit with a goldthin film patterned meander line inductor and interdigitated capacitiveelectrodes form a passive wireless telemetry system that eliminates theneed for onboard power sources and any external connections. The thinfilm sensing elements on silk platforms can be bioresorbed andintimately integrated on to biological tissues, bones or teeth. Thenanometer size scale and large surface area of the sensing elementsallows it to conform to the curvilinear surfaces biological tissues orbone. Gradual bioresorption of the supporting silk films leaves theultra-thin sensors intimately in place allowing sensitive detection oftarget analytes and subsequent wireless read-out.

Silk thin films serve as the preferred temporary platform for thesensing elements due to their optical transparency, mechanicalrobustness, bioresorbability, flexibility and biocompatibility. Whencrystallized in air, silk fibroin secondary structure kinetically favorssilk I formation, a disordered collection of α-helices and random coilsresulting in aqueous solubility. Silk I thin films are flexible,biocompatible, and possess programmable solubility rates dependent onexternally induced β-sheet content and fibroin concentration, makingthem ideal substrates for the clean transfer of graphene to biologicaland material surfaces.

Functionalization of the graphene nanosensors with bifunctional peptidesconsisting of biocombinatorially derived graphene binding motif linkedwith robust and naturally occurring antimicrobial peptide basedbiorecognition moieties enables selective and efficient recognition ofpathogenic bacteria. Non-covalent modification of graphene withbiocombinatorially derived peptides through rigorous phage displayscreening provides a general approach for the selectivefunctionalization of graphene without modifying its excellentproperties. Assembling of the designer bifunctional peptides withspecific binding domains offers a simple and versatile means tointegrate the smart property of specificity in biological recognitionpossessed biological macromolecules with the highly sensitive signaltransduction capability of graphene nanosensors. The isolation ofgraphene-binding peptides (GBP) through phage display reveals highsurface coverage and strong binding activity, which occurs throughp-stacking interactions between aromatic residues. Graphene sheetsfunctionalized with Odorranin-HP, an antimicrobial peptide isolated fromthe skin of the frog species Odorrana grahami, enable simultaneousdetection of Gram-positive and Gram-negative bacteria species. Bacterialbinding of AMPs are observed as precursor to their bacteriocidalactivity. O-HP in particular shows potent activity against Helicobacterpylori (MIC: 20 ug/mL), a Gram-negative species found in the stomach andsaliva which is estimated to be responsible for the development of over90% of duodenal ulcers and stomach cancers; Staphylococcus aureus (MIC:5 ug/mL), a Gram-positive species found on skin flora and hospitalequipments which is notoriously drug-resistant and responsible for over500,000 post-surgical wound infections in the US per year; andEscherichia coli (MIC: 30 ug/mL), a Gram-negative species found in thelower intestine of endotherms with known strains capable of acute foodpoisoning and urinary tract infections resulting from unhygienic meatand dairy preparation. O-HP is also known to exert antimicrobialactivity against methicillin resistant strains of S. aureus.

A single layer thin film inductor-capacitor (LC) resonant circuitintegrated in parallel combination with the resistive graphene monolayerenables wireless read-out and battery-free operation. The change inconductance of the graphene nanosensor up on the binding of pathogenicbacteria to the immobilized AMPs is resolved from change in thecharacteristic frequencies and bandwidth of sensor resonance.

The characteristic frequencies and the bandwidth are quantities that areinherent to the resonant circuit and do not depend on the mutualinductance (coupling coefficient) between the sensor and the readercoil. Therefore the relative alignment and location of the biosensorwith respect to the reader antenna is not important during measurements.

In one embodiment of the biosensor, the direct integration of the highlysensitive graphene nanosensors with human body and other analytesubstrates of importance such as an IV bag, to function as a fullystandalone, battery free biological sensors for the remote monitoring ofpathogenic bacteria and other biothreat agents fulfills a long-desiredneed in bio-and chemical analysis. It is contemplated the demonstratedbiosensor can be combined with other known biological binding proteins,peptides and motifs to provide biological and chemical detection systemsfor applications including, diagnostics, hospital sanitation monitoringand food safety analysis.

In another embodiment of the nanoscale form of the biosensor thebiosensor is comprised of as few components as bare graphene fashionedinto a sensor combine with a passive wireless transmitter printed on abioresorbable substrate. In this form the biosensor can be situated onnearly any biologic substrate, or any other substrate capable ofbioresorption, for a non-limiting example: cheese. This particularembodiment is not limited to combination with antimicrobial peptides,but can be used with an detection molecule capable of creating avariance in impedance including but not limited to antibodies, DNAfragments, RNA fragments, peptide binding motifs, and polymers. Inperhaps the most basic for of the nanoscale embodiment of the biosensorit has been demonstrated that the biosensor featuring only a wirelesspassive transmitter and a bare microarray without any specific detectionmolecules attached is able to detect human breath in real time.

It is then further contemplated that in certain of these types ofembodiments of the biosensor the biologic binding molecule is a glucosetransporter (for non-limiting examples GLUT1, GLUT2, GLUT3, and GLUT4)or a glucose transporter glucose-binding motif (for non-limitingexamples the glucose-binding motif of GLUT1, GLUT2, GLUT3, and GLUT4).In these embodiments the biosensor is capable of detecting the presenceand concentration of glucose in a sample, for instance a blood sample.It is yet further contemplated that certain types of these embodimentsthe biosensor could be injected intravenously, implanted intravenouslyor implanted on a body tissue in contact with blood inside a body. Inthese particular embodiments, the biosensor could be able to bindglucose in the blood and transmit the variance of impedance to anexternal receiver that could calculate the concentration of glucose inthe blood and display this concentration on the display unit.

It is also contemplated that other certain types of these embodiments ofbiosensor can be constructed with multiple circuits each comprising amicroarray, a wireless telemetry unit and a binding molecule laid on asingle substrate of a bioresorbable material. In these embodiments thebiosensor can comprise a plurality circuits with different bindingmolecules. For non-limiting instance, it is contemplated that oneembodiment of the biosensor may comprise a plurality of circuitscomprising a binding molecule that binds one of following targets: E.Coli bacteria, Staphylococcus bacteria, Rh factor,

bile, meconium, markers of meconium, herpes bacteria and, optionally, acircuit that comprises microarray, a wireless telemetry unit and acarbon microphone, a capsule containing carbon granule situated betweento metal plates, electrically coupled to the microarray. This embodimentwhen implanted in the vagina or cervical canal of a pregnant female thebiosensor acts as an internal fetal monitor, monitoring for a variety ofpotential fetal infections, complications and optionally heartbeat.

Therefore, it is contemplated that the nanoscale embodiment of thebiosensor can be combined with virtually any known sensing moleculeincluding but not limited to, antimicrobial peptides, antibodies,modified antimicrobial peptides, modified antibodies, chimeric peptidescontaining antimicrobial peptide binding motifs, chimeric peptidescontaining antibody binding motifs, DNA fragments, RNA fragments,peptide binding motifs, proteins, small molecules and polymerscombinations thereof.

EXAMPLES Example 1

Antimicrobial Peptides and Bacterial Cells. Antimicrobial peptidemagainin I (GIGKFLHSAGKFGKAFVGEIMKS) (SEQ ID NO: 1), chemicallysynthesized to contain a C-terminal cysteine residue via standardN-fluorenylmethoxycarbonyl (FMOC) solid phase peptide synthesis, wasobtained from Anaspec (San Jose, Calif.). Magainin I was also customsynthesized with an N-terminal Cysteine to compare the bacterial bindingactivity. Heat-killed pathogenic bacterial cells of E. coli O157:H7,Salmonella typhimurium and Listeria monocytogenes were purchased fromKPL (Gaithersburg, Md.). Heat-killed cells of a non-pathogenic strain ofE-Coli (ATCC 35218) was obtained from American Type Cell Culture (ATCC,Manassas, Va.) for a control. The stock solution of AMP was prepared bythe reconstitution of the lyophilized product in phosphate bufferedsaline (Sigma-Aldrich, St. Louis, Mo.) consisting of 137 mM NaCl, 2.7 mMKCl, 4.4 mM Na₂HPO₄ and 1.4 mM KH₂PO₄ (pH 7.4). The heat-killedbacterial cells were rehydrated in PBS, according to ATCCrecommendations.

The bifunctional peptide GBP-GGG-OHP (HSSYWYAFNNKT-GGGGLLRASSVWGRKYYVDLAGCAKA) (SEQ ID NO: 20) containing a graphene bindingmotif linked to the antimicrobial peptide OHP through a triglycinelinker were chemically synthesized and obtained from Peptide 2.0 Inc.,(Chantilly, Va.). The peptides were also custom synthesized with theantimicrobial peptide linked in an inverted fashion to compare thebacterial binding activity. Heat-killed pathogenic Gram-negativebacterial cells of E. coli O157:H7 and H. pylori were purchased from KPL(Gaithersburg, Md.). Heat-killed Gram-positive bacterial cells of S.Aureus were purchased from Invitrogen Inc. The stock solution of peptidewas prepared by the reconstitution of the lyophilized powder in DIwater. Different concentrations of bacterial samples were prepared fromstock solutions by dilutions in deionized water. Phosphate bufferedsaline consisting of 137 mM NaCl, 2.7 mM KCl, 4.4 mM Na₂HPO₄ and 1.4 mMKH₂PO₄ (pH 7.4), FeCl₃, Sodium carbonate and Lithium bromide for theprocessing of silk is purchased from Sigma-Aldrich (St. Louis, Mo).

Example 2 Interdigitated Microelectrode Array (IMA) and MicrofluidicFlow Cell

Interdigitated capacitive electrodes were microfabricated on 4″ p-typesilicon wafers (boron-doped, <100>, 10-16 Ω-cm, 550 μm thick). A 1 μmthick silicon dioxide layer was deposited on the wafer by plasmaenhanced chemical vapor deposition (PECVD) to form electrical insulationbetween the Si substrate and the capacitive electrodes. S1813photoresist was patterned using photolithography, followed byelectron-beam evaporation of 10 nm Ti and 300 nm Au under high vacuumconditions. The titanium layer helps to provide adhesion of the Au tothe SiO₂. The IMA was then finally developed by lift-off patterning ofthe metallic layer in acetone with ultrasonic activation. The electrodearray consisted of 50 pairs of interdigitated capacitive electrodes withan electrode width and separation of 5 μm. A polydimethylsiloxane (PDMS)microfluidic flow cell consisting of a detection microchamber with anembedded microelectrode array, inlet and outlet ports was formed bybonding the IMA chip to the PDMS channel. The PDMS micro-channel formedon the master mold was partially cured, aligned with the microelectrodearray and bonded by permanently curing at 80° C. for 2-3 hr.Microfluidic connectors were fixed on to the inlet and outlet portsthrough drilled holes.

Example 3 Sensor Surface Functionalization with Magainin

The immobilization of the peptides to a gold surface was performed atechnique utilizing the native thiol groups present in cysteineresidues. Furthermore, cysteine residues can be synthetically introducedin to a specific site of the peptide to form a properly orientedrecognition layer. Magainin I was synthesized both with a C-terminal andN-terminal cysteine via FMOC solid-phase synthesis. Prior to theimmobilization procedure, the gold IMA electrodes were cleansed usingacetone, isopropanol and deionized water. Stock solutions of the AMPswere prepared in phosphate-buffered saline (PBS), pH 7.4, consisting of137 mM NaCl, 2.7 mM KCl, 4.4 mM Na₂HPO₄ and 1.4 mM KH₂PO₄. For theimmobilization of the AMPs, 800 μg/ml (unless otherwise mentioned) ofmagainin I in PBS buffer was injected into the sensing chamber andincubated for 60 min under static conditions. The functionalizedelectrodes were then rigorously washed with 1×PBS to remove any unboundAMPs, rinsed with de-ionized water and dried in liquid nitrogen.

Example 4 Fluorescent Microscopy

Stock solutions of propidium iodide (PI), nucleic acid stain (Molecularprobes, Eugene, Oreg.) were made from solid form by dissolving PI(MW=668.4) in deionized water at 1 mg/mL (1.5 mM) and stored at 4° C.,protected from light. Heat-killed bacterial cells rehydrated in PBS werethen incubated with 3 μM solution of propidium iodide, PI (made bydiluting the 1 mg/mL stock solution 1:500 in buffer) for 10-15 min (51).After incubation, the cells were centrifuged into pellets, thesupernatant was removed and the cells were re-suspended in fresh 1×PBSbuffer. The samples of stained bacterial cells (E. coli O157: H7,Salmonella, non-pathogenic E. coli and Listeria, all 10⁷ CFU/mL) werethen allowed to incubate with the immobilized Magainin for 15-20 min inthe dark. After incubation, the Au surfaces were washed with PBS bufferand dried under liquid nitrogen. The binding pattern of the differentbacterial cells was imaged using a Zeiss axiovert inverted microscopeand recorded with a Zeiss axiocam digital camera. Surface density of thebound bacterial cells was analyzed and plotted using imageJ softwarefrom NIH.

Example 5 Impedance Spectroscopy

Dielectric property changes due to AMP-bacteria interactions were probedusing a Fast-Fourier Transform (FFT) spectrum analyzer. The dielectricproperties were investigated over a frequency range of 10 Hz to 100 kHz,with 0 V DC bias and 50 mV AC signals using a SRS 785, 2-channel dynamicsignal analyzer. The LabView program routine was used to collect andrecord the data through a GPIB interface. An external op-amp amplifiercircuit was used to minimize the noise and a MATLAB program was used toplot the impedance spectra from the analyzer output (see supplementaryinformation FIG. 51). For sensitivity measurements, pathogenicGram-negative E. coli O157:H7 bacterial cells were injected into themicrofluidic flow channel at various dilutions, and incubated with theimmobilized magainins for 12-15 min, under static conditions. To ensurethe response of the sensor toward bound bacterial cells, the impedancespectrum was taken after the removal of unbound cells by thoroughwashing in PBS. For real-time measurements, the impedance vs. time datawas recorded while buffer solutions or different dilutions of bacterialsolutions flowed through the microfluidic channel. The flow cell wasfirst flushed with 1× PBS buffer at a flow rate of 100 μL/min toestablish a baseline. Bacterial detection measurements were performedwith the sample flowing at a rate of 5 μL/min. The main effect of thebacterial cells bound to the immobilized magainins on the impedancesignal is due to the dielectric property of the cell membrane.

Example 6 Silk Films

Silk solutions are versatile and can be prepared in various formsincluding electrospun fibers, films, gels and sponges, all havingtunable properties such as solubility, elasticity and biodegradability.Silk from the domesticated silkworm, Bombyx mori, is perhaps the mostuseful type due to its availability. B. mori cocoons are primarilycomposed of two types of protein: fibroin and sericin. The fibroinfibers, consisting of a heavy and light chain connected by a disulfidebond, are responsible for the structural integrity of the cocoons. Thefibroin protein consists of layers of antiparallel beta sheets, largelyresponsible for the tensile strength of the material. Its primarystructure mainly consists of the recurrent amino acid sequence(Gly-Ser-Gly-Ala-Gly-Ala)n (SEQ ID NO: 21). The fibers are enveloped bythe hydrophilic protein sericin, which acts as a sticky glue holding thecocoon together. An outer coating provides extra protection frommicrobes and predators.

In order to extract the desired fibroin protein, the sericin was removedby boiling cleaned cocoons in 0.02 M Na₂CO₃ for 30 minutes followed bythoroughly rinsing with water. The degummed silk was then dissolved in9.3M aqueous lithium bromide and the solution is dialyzed to removeexcess salt. The resulting aqueous solutions are 6-10% (w/v) fibroin andcan be preserved by storage at 5° C. Silk films were made by castingfibroin solutions onto PDMS and drying in air for 12-24 h, depending onthe thickness. When crystallized in air, silk fibroin secondarystructure kinetically favors silk I formation, a disordered collectionof a-helices and random coils resulting in aqueous solubility.

Silk I thin films are flexible, biocompatible, and possess programmablesolubility rates dependent on externally induced b-sheet content andfibroin concentration, making them ideal substrates for the cleantransfer of graphene to biological and material surfaces.

Example 7 Fabrication of Silk/Graphene Wireless Sensing elements.

Bioresorbable, thin film sensing elements can be patterned on to watersoluble silk fibroin films of ca 50 microns thick through a series ofsimple fabrication steps. First, dissolvable films of silk were castdirectly from aqueous Bornbyx mori (dissolved cocoons of silk worms)silk solutions to serve as bioresorbable substrates. CVD grown graphenemonlolayers from Ni thin films were released and transferred on to PDMSstamps by the removal of Ni layer in FeCl₃ and are then transfer printedon to the silk films. Clean transfer of graphene monolayers on to silkwas achieved through controlled moistening of the silk surface. Planarinductive and capacitive elements were then incorporated on to thegraphene/silk samples to enable wireless interrogation capability. Ameander line design for the inductive element allowed planarimplementation of the wireless telemetry unit, integrated on to thegraphene nanosensor via a shadow mask assisted one step evaporationelectron beam evaporation of gold. The thin film sensing elements onsilk platforms were then bioresorbed and intimately integrated on tosurfaces such as biological tissues, bones or teeth.

Example 8 Bioresorption

Integration of the wireless sensing elements on to tooth surface wasachieved through the dissolution of the supporting silk substrate. Thesurface of the tooth was first cleaned and dried with a low-particulateclean room wipe. In the case of teeth, and other such dry surfaces, amoistened cotton swab was used to slightly wet the surface. Thegraphene-Au electrode sensing elements on the temporary silk films werethen carefully aligned and placed on the tooth surface with the deviceside facing down. The temporary silk film platform was then graduallydissolved off by the application of water, leaving the ultra-thinsensors intimately attached to the surface of the tooth. The integrationof the sensing elements on to the surface of the I.V. bag was also donesimilar to the above procedure. In the case of wet surfaces ofbiological tissue and food materials such as cheese, the dissolution ofthe silk film was observed to be faster. The high surface area of thegraphene and electrodes ensures high adhesive conformability to thecurvilinear surfaces in biological tissues such as skin or tooth.

Example 9 Graphene Functionalization with Antimicrobial Peptide O-HP

In functionalizing graphene sheets, a peptide that links GBP and O-HPvia triglycine was generated, resulting in the 38 amino acid sequence,HSSYWYAFNNKT-GGG-GLLRASSVWGRKYYVDLAGCAKA (SEQ ID NO: 20): TheGBP-GGG-O-HP bifunctional peptide. These peptides were dissolved indeionized water at a concentration of 1 mg/mL. Five microliters of thepeptide solutions were dropped onto the graphene and incubated for 15min, followed by washing with deionized water and drying.

Example 10 Single Bacterium Detection Measurements

Electrical measurements for the detection of single bacteria bindingmeasurements were performed with a lock-in detection system usingStanford Research Systems 810 DSP lock-in amplifier. A modulation signalof 50 mV was used with a modulation frequency of 73 Hz with zero DC biasto avoid any electrochemical reactions. The resistance of the graphenesensor was monitored continuously in the presence of analyte sample ofvarious dilutions of bacterial cells.

Bacterial cells for the simultaneous optical monitoring and for theantimicrobial peptide-bacteria binding density studies are fluorescentlylabeled with propidium iodide using known methods. Stock solutions ofpropidium iodide (PI), nucleic acid stain (Molecular probes, Eugene,Oreg.) were made from solid form by dissolving PI (MW=668.4) indeionized water at 1 mg/mL (1.5 mM) and stored at 4° C., protected fromlight. Heat-killed bacterial cells rehydrated in PBS were then incubatedwith 3 μM solution of propidium iodide, PI (made by diluting the 1 mg/mLstock solution 1:500 in buffer) for 10-15 min. After incubation, thecells were centrifuged into pellets, the supernatant was removed and thecells were re-suspended in DI water or 1×PBS buffer. The binding patternof the different bacterial cells was imaged using a Zeiss axiovertinverted microscope and recorded with a Zeiss axiocam digital camera.For the real-time detection of single bacterial cell (E. coli O157: 117)a sample containing100 bacterial cells per μL was loaded by pipette onto the graphene sensors. Simultaneous bright field and fluorescent datawere recorded together with lock-in resistance data with the focal planeset on the graphene surface to identify the events when the bacterialcells come close to the sensor. The motion of the bacterial cells weretracked relative to the position of the graphene surface with the helpof video spot tracker, a freely available automated tracking softwareand also with the manual tracking plugin in the National Institutes ofHealth's ImageJ software.

Example 11 Wireless Sensing

An inductive-capacitive (LC) resonant circuit, integrated in parallelwith the resistive graphene monolayer, forms the basis of the wirelessread-out unit. The reader device consists of a two-turn coil antennaconnected to a frequency response analyzer (HP 4191A RF impedanceanalyzer). The wireless reader, which is powered by an alternatingcurrent source, is responsible for wirelessly transmitting power andreceiving sensor data from the remote passive sensor, all throughinductive coupling. Passing an AC signal through the antenna generates amagnetic field, e.g., reference number 28, FIG. 11C, inducing currentvia mutual inductance in the coil of the sensing element (Faraday'slaw), finally resulting in a potential drop that depends on theconductance of the graphene nanosensor. Any change in the conductivityof the sensor system resulting from biological or chemical changeshappening at the transducer surface will be reflected as a change in thefundamental properties—characteristic frequencies and the bandwidth—atresonance. This allows the reader to wirelessly interrogate the sensingelement via the complex impedance of the system. The equivalent circuitof the sensing element can be used to calculate the bandwidth of thereader-sensor system and the input impedance of the system as viewed bythe reader device. The change in the capacitance of thegraphene-interdigitated electrode sensing system is deduced from theresonance frequency shift, the expression for which is as follows:

$\begin{matrix}{{\Delta \; f} = \frac{1}{2\; \pi \sqrt{L*\Delta \; C}}} & (2)\end{matrix}$

The bandwidth of the sensor resonance is measured from the resonancepeaks and was used to calculate the change in resistance (conductance)of the graphene-electrode sensing element up on the binding of thebacterial cells. The expression for the change in resistance of thesystem is as follows:

ΔR=1/(2π*ΔC*ΔB)   (3)

Detailed description of the calculation of the system bandwidth isprovided in the supporting information. The characteristic frequenciesof the sensor system can be measured by monitoring the complex impedancespectrum of the system. The frequency at which the imaginary part of thesensor impedance vanishes (reactance X=0) is continually monitored andused to calculate the change in the graphene conductance.

$\begin{matrix}{{\Delta \; R} = \left( \frac{L}{\Delta \; {C\left( {1 - {2\pi*\Delta \; f_{z}}} \right)}L*\Delta \; C} \right)^{1/2}} & (4)\end{matrix}$

Although features and elements are described above in particularcombinations, each feature or element may be used alone without theother features and elements or in various combinations with or withoutother features and elements.

What is claimed is:
 1. A biosensor configured for detection of a targetand for application to a test surface, the biosensor comprising: amicroelectrode array having an electrical conductivity that variesdepending on a target concentration, the microelectrode array beingdisposed on a dissolvable substrate, the microelectrode array beingconfigured to generate a signal corresponding to the electricalconductivity of the microelectrode array; and the dissolvable substratebeing configured to be dissolved off by the application of water,leaving the microelectrode array intimately attached to the testsurface.
 2. The biosensor of claim 1, further comprising a transmitterelectrically coupled to the microelectrode array and forming an electriccircuit, the transmitter being configured to transmit the signalcorresponding to the electrical conductivity of the microelectrodearray, the dissolvable substrate being configured to be dissolved off bythe application of water, leaving the microelectrode array and thetransmitter intimately attached to the test surface.
 3. The biosensor ofclaim 1, further comprising a binding molecule disposed on themicroelectrode array.
 4. The biosensor of claim 3 wherein the bindingmolecule is configured to interact with and attract a subsequent bindingmolecule, the microelectrode array having an electrical conductivitythat is dependent on a binding state of the binding molecule.
 5. Thebiosensor of claim 3, wherein binding molecule is selected from the listconsisting of: antimicrobial peptides, antibodies, modifiedantimicrobial peptides, modified antibodies, chimeric peptidescontaining antimicrobial peptide binding motifs, chimeric peptidescontaining antibody binding motifs, DNA fragments, RNA fragments,peptide binding motifs, proteins, small molecules and polymerscombinations thereof.
 6. The biosensor of claim 1, wherein thedissolvable substrate is selected from a list consisting of: cellulose,paper, silk, silk fibroin and combinations thereof.
 7. The biosensor ofclaim 1, wherein the dissolvable substrate is placed in contact with abiologic test surface.
 8. The biosensor of claim 7, wherein the biologictest surface is capable of bioresorption.
 9. The biosensor of claim 7,wherein the biologic test surface is selected from the list consistingof: teeth, bone, skin, tissue, hair, nail, cornea, gum, tongue, palate,brain, heart, lung, membrane, leaf, root, bark, fur, feather, chiton andscale.
 10. The biosensor of claim 2, wherein the transmitter is awireless transmitter.
 11. The biosensor of claim 10, further comprisinga receiver unit having a planar coil antenna electrically coupled to themicroelectrode array, the planar coil antenna being configured to powerthe wireless transmitter, the wireless transmitter being configured tosend a wireless signal corresponding to the electrical conductivity ofthe microelectrode array.
 12. The biosensor of claim 1, wherein thebinding molecule is contains a glucose-binding motif.
 13. The biosensorof claim 1, wherein the microelectrode array is interdigitated.
 14. Thebiosensor of claim 1, wherein the microelectrode array comprisesgraphene.